Device and method for determining depth and concentration of a subsurface fluorescent object

ABSTRACT

A method and device for determining the depth and fluorophore concentration of a fluorophore concentration below the surface of an optically absorbing and scattering medium suitable for use in fluorescence-based surgical guidance such as in tumor resection is described. Long-wavelength stimulus light us used to obtain deep tissue penetration. Recovery of depth is performed by fitting measured modulation amplitudes for each spatial frequency to precomputed modulation amplitudes in a table of modulation amplitudes indexed by optical parameters and depth.

PRIORITY CLAIM

The present application is a continuation of U.S. patent applicationSer. No. 17/050,745, filed Oct. 26, 2020, which is a 35 U.S.C. § 371filing of International Application No. PCT/US2019/029421 filed 26 Apr.2019, which claims priority to U.S. Provisional Patent Application No.62/663,158 filed 26 Apr. 2018, each of which is herein incorporated byreference in its entirety.

GOVERNMENT RIGHTS

This invention was made with government support under grant no. R01NS052274 awarded by the National Institutes of Health. The governmenthas certain rights in the invention.

TECHNICAL FIELD

A method and device for determining depth and fluorophore concentrationof fluorescent objects located below a surface of an optically absorbingand scattering medium is described. The method and device are suitablefor use in fluorescence-based surgical guidance such as in tumorresection.

BACKGROUND

Surgery is used commonly as part of the treatment of patients withtumors. In many cases, the goal is to maximize the extent of tumorresection, while minimizing damage to surrounding heathy tissues toimprove survival and quality of life. Preoperative radiological imagingsuch as magnetic resonance imaging (MRI) or X-ray computed tomography(CT) is often used to guide the procedure, for example to show thelocation of the tumor with reference to other anatomical structures.However, during surgery, tumors may shift in position, leading to lossof co-registration with pre-surgical images. In addition, it may bedifficult to identify the true tumor margin. Near the end of surgery, itmay also not be possible for the surgeon to visualize residual tumorthat could regrow if not removed.

Hence, fluorescence imaging (fluorescence guided resection, FGR) duringsurgery is emerging as a valuable adjunctive technique to address theselimitations.

FGR may be based on the natural fluorescence of the tissue, often calledtissue autofluorescence. Alternatively, an exogenous fluorescentmaterial (“fluorophore”) may be administered to the patient to highlightthe tumor. In this case the spectrum of fluorescent light emerging fromthe tissue upon illuminating the tissue with light of a suitableexcitation wavelength is the combination of tissue autofluorescence plusfluorescence from the exogenous fluorophore at the concentration that itreaches in the particular tissue.

Different exogenous fluorophores have been investigated both inpreclinical models and in clinical trials of different tumor locationsand stages. For example, indocyanine green (ICG), that emitsfluorescence in the near-infrared region of the spectrum is often usedto image the perfusion of blood within tissue. For tumor imaging, and inparticular for guiding surgery, a common fluorophore is protoporphyrinIX (PpIX) that is synthesized within cells upon administration ofaminolevulinic acid (ALA). In general, this shows high contrast betweentumor and normal host tissue. The PpIX absorbs at a range ofwavelengths, from around 400 nm to around 630 nm, exciting a spectrum offluorescence emission at around 700 nm. Absorption of PpIX is farstronger at the short-wavelength end of this absorption range than atthe long end, as illustrated in FIG. 1B. Because of greater absorptionefficiency and easier filtering of emissions from excitation light atwide separations, systems imaging ALA-induced PpIX fluorophoreconcentrations typically use short-wavelength, or “blue,” light forexcitation while observing deep red emissions light.

ALA-PpIX fluorescence image guidance is typically used, for example inbrain tumor surgery, by incorporating a blue light source for excitationand red-to-near infrared fluorescence detection into a neurosurgicalmicroscope. For other sites, the fluorescence imaging technology isincorporated into an endoscope or is “stand-alone”, independent of othermedical imaging devices. ALA-PpIX fluorescence image guidance hasdemonstrated clinical benefit by enabling more complete resection of thetumor near the end of surgery. For example, ALA-PpIX-based FGR ofhigh-grade glioma using blue-light excitation has been shown to at leastdouble the rate of complete resection and improve progression-freesurvival.

When fluorescence images are taken at a tissue surface, the strength ofthe fluorescence signal is affected by the attenuation of the excitationand emission light due to the optical absorption and scattering of thelight by the tissue. This attenuation can be substantial and shows acomplex wavelength dependence. As a result, a relatively weak measuredfluorescence signal or low-brightness image taken at the tissue surfacemay be the result of either low concentration of the fluorophore or highabsorption and/or scattering at the excitation and/or emissionwavelengths. Conversely, a region of bright fluorescence intensity maybe due to high fluorophore concentration and low absorption andscattering. The absolute concentration of the fluorophore, for a givenadministered dose for the fluorophore (or its precursor as in the caseof ALA-PpIX) is an important marker of the presence of malignant cellsor tissues so that resolving such ambiguities is of value.

Thus, measuring the absolute PpIX concentration in the tissue duringresection can significantly improve the sensitivity and specificity ofresidual tumor detection, by better identifying fluorescent tissue thatis not seen by qualitative visual assessment. This has been demonstratedusing a fiberoptic probe that measures the fluorescence spectrum oflight, at a single specific location the tissue surface. The spectrum ofdiffusely-reflected light is measured and analyzed to calculate theabsorption coefficients (μ_(a)) and the transport scatteringcoefficients (μ′_(s)) of the tissue at the fluorescence excitation anddetection wavelengths. After applying appropriate calibration factors,these data are then used as input to a model of light transport intissue to convert the measured fluorescence signal into the true, orintrinsic, signal corrected for the confounding effects of unknown lightattenuation by the tissue. The PpIX concentration in the tissue at theprobe location is then recovered by spectrally un-mixing the known PpIXfluorescence spectrum from the tissue autofluorescence and from anyphotoproducts of PpIX that may have been generated during the surgery.

In this point-spectroscopic mode, quantitation of the PpIX concentrationin the tissue has enabled enhanced resection rates in both low- andhigh-grade gliomas over and above the rates achieved using quantitativefluorescence image guidance alone. This technique is referred to here asquantitative point fluorescence spectroscopy, qFS.

SUMMARY

We will use ALA-PpIX fluorescence guided resection of brain tumorresection to demonstrate the present invention but it is understood thatthe method and device described herein are applicable to otherfluorophores, including near-infrared fluorophores, and for differenttumors and non-oncological applications.

Quantitative localized fluorescence spectroscopy has been extended towide-field quantitative imaging, qFI, in which μ_(a) and μ′_(s) of thetissue are first mapped across the surgical field-of-view and then thesedata are applied to correct the measured fluorescence images on apixel-by-pixel basis, analogously to the point-probe technique. Onemethod to map the optical properties is by spatial frequency domainimaging (SFDI). This is a specific form of diffuse optical imaging thatcan produce images of μ_(a) and μ′_(s) parameters over large fields ofview with submillimeter spatial resolution. In SFDI, sine-wave or otherperiodic patterns of light are projected onto the tissue surface, andthe pattern of spatially-resolved diffuse reflectance from the tissue isimaged. This pattern has the same spatial frequency and phase as theprojected pattern but intensity distribution is modulated by scatteringand attenuation in the tissue.

Knowing these optical absorption and transport scattering maps acrossthe surgical field of view, the effects of light attenuation on thedetected fluorescence image are corrected for, analogously to qFS. Aftersubtracting the tissue autofluorescence, an image of the absoluteconcentration of the PpIX at or near the tissue surface is generated. Adevice in which SFDI is incorporated into a hyperspectral fluorescenceimaging system has been reported and the PpIX concentration inintracranial tumors in a glioma model in rats was calculated in the formof 2D maps, with an accuracy of about ±˜10%, which is comparable to thatof qFS. The minimum detectible concentration was about 13 ng/ml.

The methods and devices for qFS and qFI as defined above that applyrigorous analytical or numerical modeling of light propagation in tissueare distinct from f other methods and devices based on semi-empiricalalgorithms. For example, fluorescence images may be collected at twodifferent excitation wavelengths or at two different detectionwavelengths, and various ratio images can be generated from these. Thesemethods and devices are intended to minimize or partly compensate forthe effects of light attenuation and other factors that may affect thefluorescence image. However, they differ from the present inventionbecause they do not enable the absolute fluorophore concentration in thetissue to be calculated and mapped.

When light at short wavelengths such as in the violet or blue region ofthe spectrum is used to excite the fluorophore, the limited penetrationdepth of this light in tissue restricts qFI to superficial tumors, up toa depth of about 1 mm. The purpose of the present invention is to extendthe quantitative capability of qFI to deeper tissues lying beyond theresection bed surface to image both the depth of sub-surface tumor andthe fluorophore concentration therein. The present invention differsfrom the above qFI methods and devices in that the fluorophoreconcentration is calculated for fluorophore lying substantially belowthe tissue surface.

BRIEF DESCRIPTION OF THE DRAWINGS

The elements in the figures are not to scale.

FIG. 1A is a schematic of a benchtop hyperspectral SFDI instrument usedon tissue-simulating phantoms and gliomas or other cancers as describedherein.

FIG. 1B is a graph illustrating absorption of protoporphyrin IX (PpIX)versus wavelength, showing that absorption of excitation wavelengthlight at short wavelengths is dramatically higher than at longwavelengths near the 632-nanometer wavelength used for excitation in thepresent system.

FIG. 2 is a block diagram of an alternative instrument for using qdFIduring surgery, such as brain surgery for removal of gliomas.

FIG. 3 is a flowchart of a method of operating the instruments of FIGS.1 and 2 to perform qdFI and present images of use in analyzingpathological specimens or of indicating tumor tissue to a surgeon sothat the surgeon may remove the tumor tissue during surgery.

FIG. 4 is a graph of the measured normalized modulation amplitude of thefluorescence signal at 705 nm as a function of the spatial frequency ofthe incident 632 nm excitation light for different depths of thefluorescent object below the surface of the liquid phantom. The depthsin mm are indicated on each curve. The fluorescent object contained 10μg/ml of PpIX mixed with blood and Intralipid to simulate tissue opticalproperties of μ_(a)=0.001 lines per millimeter (mm⁻¹) and μ′_(s)=1 mm⁻¹at 705 nm. Error bars are the standard deviation across threereplicates.

FIG. 5A-5C shows representative plots of the fluorescence modulationversus inclusion depth for different combinations of optical absorptionand transport scattering coefficients of the phantom at 705 nm and for aparticular spatial frequency f.

FIG. 6A-6C shows plots of the calculated depth versus the actual depthof the fluorescent inclusion for phantoms of different absorption andtransport scattering coefficients. Error bars are the standard deviationfrom 3 replicates. The line of best fit is also shown together with theR² goodness of fit.

FIG. 7A-7C shows plots of the calculated PpIX concentration versus theactual depth of the fluorescent inclusion for phantoms of differentabsorption and transport scattering coefficients. The true value of thePpIX concentration is indicated by the horizontal line in each plot.Error bars are the standard deviation from 3 replicates. The p valuesindicate whether or not the calculated concentrations are statisticallydifferent from the actual concentrations.

DETAILED DESCRIPTION OF THE EMBODIMENTS AND SUPPORTING DATA

There are at least two complementary approaches to determine the depthof subsurface fluorescent tumor, in a form of imaging referred to hereas dFI. In one method and device, fluorescence images are taken atmultiple different excitation wavelengths. The images are then rankedaccording to the expected effective penetration depth of each wavelengthof light in the tissue, a form of intensity thresholding is applied, anda topographic image is generated that displays the depth of the toplayer of the subsurface fluorescent tumor. For PpIX this method anddevice determined the depth of a subsurface fluorescent object with anaccuracy of about ±0.5 mm up to a depth of about 3 mm in brain tissue.

The second method and device, which has also been demonstrated for PpIX,is to excite the fluorescence with red light that penetrates much deeperinto the tissue, and then to form images at several differentwavelengths of the fluorescence emission. Since longer wavelengths areless attenuated by the overlying tissue, the detected fluorescencespectrum at each point in the image is distorted to a degree thatdepends on the depth from which the fluorescent light originates. ForPpIX this method and device have given a depth accuracy of about ±1 mmup to a depth of about 8 mm in brain tissue. Red light is lightgenerally of 630 nanometers or longer wavelength.

The present invention differs from these two methods and devices for dFIin that the subsurface fluorophore depth images are here calculated froma series of SFDI images taken in fluorescence mode under long wavelengthexcitation (e.g. red light in the case of PpIX) The depth estimate isthen based on the rate of decay of the fluorescence-mode SFDI signalmodulation with increasing spatial frequency.

Qualitative ALA-PpIX FGR using red light to excite PpIX has beenreported in 30 patients with various intracranial tumor pathologies,adapting a fluorescence-enabled neurosurgical microscope to thisexcitation wavelength. This has demonstrated enhanced tumor resection byenabling detection of sub-surface PpIX fluorescence up to an estimateddepth of about 5 mm, compared to about 1 mm under blue-light excitation.However, the PpIX absorption is 1-2 orders of magnitude lower near 635nm than at 405 nm and, due to the much longer optical paths at around635 nm, the detected fluorescence intensity depends strongly on thefluorophore concentration, depth and distribution and is also affectedby the tissue absorption and scattering. The negative impact of theseuncontrolled factors on quantifying the fluorophore concentration insubsurface tumor motivates the development and validation of methods anddevices to separate the depth of the fluorophore from its concentrationat that depth, as in the present invention.

Hyperspectral fluorescence imaging can also be used to estimatesub-surface PpIX depth, using the fact that the logarithmic ratio of twofluorescence images at different emission wavelengths is approximatelylinear with fluorophore depth and with optical penetration depth. Tumordepths from about 1 to 9 mm could be recovered within ±1 mm. However,additional constraints were needed to estimate the fluorophoreconcentration at depth. Fluorescence-mode SFDI improves spatiallocalization of the fluorophore by gating the penetration depth of theexcitation light according to spatial frequency, thereby facilitatingdepth-resolved imaging. This is to be contrasted with planarillumination, i.e. illumination of the tissue with uniform unpatternlight, where the collected signal comprises fluorescence originatingfrom all depths but weighted towards the surface.

SFDI-enabled depth-resolved fluorescence imaging has been reported.Multiple spatial-frequencies of red-light SFDI were used to excite PpIXin sub-surface intracranial tumors in mice. A 3D tomographic approachwas applied to recover the PpIX distribution and concentration. Thisdiffers from the present invention, since the tumor depth had to beknown a priori for this reported method, whereas in the presentinvention this depth is calculated from the measured data in anindividual patient without prior knowledge of the sub-surface depth ofthe fluorophore.

The present method and device determine the sub-surface fluorophoredepth by exploiting SFDI's depth encoding capabilities from just belowthe surface up to about 9 mm with red-light excitation of PpIX at around635 nm. In the present invention the images of both the depth and thefluorophore concentration at that depth for tumor lying below theaccessible tissue surface (e.g. the resection bed) are determined. Thegoal here is not to produce a full 3D tomographic distribution of thefluorophore in the tissue, but rather to provide topographic imagesindicating the depth of fluorescent tumor closest to the resectioncavity surface, as well as the fluorophore concentration at that depth.Both separately and in combination, this information should help thesurgeon make a more informed decision on whether or not to continueresection in a particular region of the surgical field. This method isreferred to here as qdFI.

To encode the fluorophore depth, patterned red light is projected ontothe sample and the spatially-modulated fluorescence images are collectedby a spectrally-resolved camera. The rate of decrease of thefluorescence modulation is calculated. The optical absorption andtransport scattering coefficient values of the tissue are determined bya separate SFDI procedure using broad-band light and imaging the diffusereflectance patterns from the tissue. These data are used, together withthis rate of decrease, to calculate the subsurface depth of thefluorescent object. The images of tissue absorption and transportscattering are then also used in a forward model of light propagation inan optically turbid medium to calculate the concentration of thefluorophore at this depth in each pixel of the image.

This disclosure therefore relates to fluorescence guided surgery, and inparticular concerns quantitative fluorescence assessment. Rather thanrelying on the qualitative visual view provided to the surgeon thetechnology estimates quantitative levels of fluorophore concentration,in this case, at depth below the immediate surgical surface. Thedisclosure includes methods to quantify the concentration of fluorophorein tissue for guiding surgery. Quantitative fluorescence imaging astaught herein is more sensitive and more accurate than relying on thesurgeon's visual perception of fluorescence. Data shows thatfluorescence which is to visible to the surgeon can represent resectabletumor in neurosurgery.

At least two complementary approaches have been presented to determinethe depth of subsurface fluorescent tumor, in a form of imaging referredto here as dFI. In one method and device, fluorescence images are takenat multiple different excitation wavelengths. The images are then rankedaccording to the expected effective penetration depth of each wavelengthof light in the tissue, a form of intensity thresholding is applied, anda topographic image is generated that displays the depth of the toplayer of the subsurface fluorescent tumor. For PpIX this method anddevice determined the depth of a subsurface fluorescent object with anaccuracy of about ±0.5 mm up to a depth of about 3 mm in brain tissue.

The second method and device, which has also been demonstrated for PpIX,is to excite the fluorescence with red light that penetrates much deeperinto the tissue, and then to form images at several differentwavelengths of the fluorescence emission. Since the longer wavelengthsare less attenuated by the overlying tissue, the detected fluorescencespectrum at each point in the image is distorted to a degree thatdepends on the depth from which the fluorescent light originates. ForPpIX this method and device have given a depth accuracy of about ±1 mmup to a depth of about 8 mm in brain tissue.

The present invention differs from these two methods and devices for dFIin that the subsurface fluorophore depth images are here calculated froma series of SFDI images taken in fluorescence mode under long wavelengthexcitation (e.g. red light in the case of PpIX) The depth estimate isthen based on the rate of decay of the fluorescence-mode SFDI signalmodulation with increasing spatial frequency.

Qualitative ALA-PpIX FGR using red light to excite PpIX has beenreported in 30 patients with various intracranial tumor pathologies,adapting a fluorescence-enabled neurosurgical microscope to thisexcitation wavelength. This work (published by the Inventors) hasdemonstrated enhanced tumor resection by enabling detection ofsub-surface PpIX fluorescence up to an estimated depth of about 5 mm,compared to about 1 mm under blue-light excitation. However, the PpIXabsorption is 1-2 orders of magnitude lower than at 405 nm and, due tothe much longer optical paths at around 635 nm, the detectedfluorescence intensity depends strongly on the fluorophoreconcentration, depth and distribution and is also affected by the tissueabsorption and scattering. The negative impact of these uncontrolledfactors on quantifying the fluorophore concentration in subsurface tumormotivates the development and validation of methods and devices toseparate the depth of the fluorophore from its concentration at thatdepth, as in the present invention.

Hyperspectral fluorescence imaging can also be used to estimatesub-surface PpIX depth, using the fact that the logarithmic ratio of twofluorescence images at different emission wavelengths is approximatelylinear with fluorophore depth and with optical penetration depth. Tumordepths from about 1 to 9 mm could be recovered within ±1 mm. However,additional constraints were needed to estimate the fluorophoreconcentration at depth. Fluorescence-mode SFDI improves spatiallocalization of the fluorophore by gating the penetration depth of theexcitation light according to spatial frequency, thereby facilitatingdepth-resolved imaging. This is to be contrasted with planarillumination, i.e. illumination of the tissue with uniform unpatternlight, where the collected signal comprises fluorescence originatingfrom all depths but weighted towards the surface.

SFDI-enabled depth-resolved fluorescence imaging has been reported.Multiple spatial-frequencies of red-light SFDI were used to excite PpIXin sub-surface intracranial tumors in mice. A 3D tomographic approachwas applied to recover the PpIX distribution and concentration. Thisdiffers from the present invention, since the tumor depth had to beknown a priori for this reported method, whereas in the presentinvention this depth is calculated from the measured data in anindividual patient without any prior knowledge of the sub-surface depthof the fluorophore.

Longwave qdFI

The present method and device determine the sub-surface fluorophoredepth by exploiting SFDI's depth encoding capabilities from just belowthe surface up to about 9 mm with red-light excitation of PpIX at around635 nm. In the present invention the images of both the depth and thefluorophore concentration at that depth for tumor lying below theaccessible tissue surface (e.g. the resection bed) are determined. Thegoal here is not to produce a full 3D tomographic distribution of thefluorophore in the tissue, but rather to provide topographic imagesindicating the depth of the fluorescent tumor closest to the resectioncavity surface, as well as the fluorophore concentration at that depth.Both separately and in combination, this information should help thesurgeon make a more informed decision on whether or not to continueresection in a particular region of the surgical field. This method isreferred to here as qdFI.

To encode the fluorophore depth, patterned red light is projected ontothe sample and the spatially-modulated fluorescence images are collectedby a spectrally-resolved camera. The rate of decrease of thefluorescence modulation is calculated. The optical absorption andtransport scattering coefficient values of the tissue are determined bya separate SFDI procedure using broad-band light and imaging the diffusereflectance patterns from the tissue. These data are used, together withthis rate of decrease, to calculate the subsurface depth of thefluorescent object. The images of tissue absorption and transportscattering are then also used in a forward model of light propagation inan optically turbid medium to calculate the concentration of thefluorophore at this depth in each pixel of the image.

We disclose the following methods and devices for qdFI. These methodsand devices are illustrated by the example of PpIX but are not limitedto this fluorophore. With suitable adjustment and knowing thefluorescence excitation and emission spectra, they may be applied toother fluorophores working across the ultraviolet-visible-near infraredranges of the optical spectrum.

Specific apparatus and experiments is described together with resultsillustrating the performance of method and device. However, the methodand device may be applied to other experimental conditions, includingfluorescence-guided surgery in patients. Neither are the method anddevice restricted to brain tumors such as gliomas. Neither are themethods and devices restricted to the application of guiding tumorsurgery, since they may be applied to determining the depth andconcentration of any subsurface fluorophore within an opticallyabsorbing and scattering medium.

An object containing a fluorescent material (the fluorophore) of knownfluorescence excitation and emission spectra is located below thesurface of a homogeneous medium of unknown optically absorbing andscattering properties. The surface of the medium is illuminated withknown patterns of spatially-modulated light, each pattern having adifferent known spatial frequency, f. Two different forms ofillumination are used one after the other in either sequence

In the first form, referred to here as fluorescence-mode SFDI, theillumination light is at a wavelength or range of wavelengths known toexcite the fluorophore and to provide sufficient penetration into themedium to interact with the fluorescent object, for example red light toexcite PpIX fluorescence in tissue, such as 632 nm wavelength red lightthat excites a fluorescence peak at 705 nm, also red light, this peaklying between 700 and 710 nanometers. The images of fluorescence at thetissue surface is collected. In the second form, referred to here asreflectance-mode SFDI, the illumination light is spectrally broad bandand images of the diffusely-reflected light from the tissue is collectedat wavelengths of interest. In alternative embodiments, otherfluorophores may be used with red or infrared excitation light and redor infrared fluorescent emissions light.

A system constructed for demonstrating qdFI, integrating hyperspectralimaging and SFDI in a single device, is illustrated in FIG. 1 . Thesystem operates to localize a sub-surface PpIX fluorescent inclusion 2in a tissue-simulating phantom 1, or a glioma or other cancer in tissue.Light from a 632 nm LED light 3 is shown as an example of one of theseveral LED sources that may be activated. This light is a coupled by alight guide 4 to a digital light projector 5 onto a mirror 6 thatreflects the spatially-modulated light pattern 7 onto the surface of thephantom 1. The spatially-modulated fluorescence images are filtered by a650 nm long-pass filter 8 and passed through relay lenses 9 and thenthrough a liquid crystal tunable filter 10 to a CCD camera 11. Thesystem includes 6 LED light sources centered at 390, 440, 475, 512, 586and 632 nm each with 20 nm bandwidth (Spectra X, Lumencor, Beaverton,Oreg., USA), of which one example is shown in the figure. Differentcentral wavelengths and bandwidths could be selected to match otherfluorophores. After passing through a liquid light guide 4 (LGG0338,ThorLabs, Montreal, Canada), the light is reflected by a spatial lightmodulator, DLP, 5 (Digital Micromirror Device, 0.55XGA Series 450, TexasInstruments, Dallas, Tex., USA) onto the surface of the medium 1. AllLEDs are turned on simultaneously to simulate white-lightreflectance-mode SFDI to map the optical absorption and transportscattering coefficients of the tissue. In fluorescence-mode SFDI onlythe 632 nm LED is activated. The intensity of each LED is selected tooptimally cover the dynamic range of the detector while avoidingsaturation. The imaging sub-system consists of a 14-bit CCD detectorarray 11 (Pixelfly, PCO AG, Kelheim, Germany) connected to avisible-range liquid crystal tunable filter 10 (LCTF: Varispec-07-02,Perkin Elmer. Inc, Waltham, Mass., USA) by a 1:1 relay C-mount lens 9,giving a field of view of 5 cm by 5 cm with 2.0 cm depth of field at thefocal plane. A 650 nm long-pass filter 8 (FELH0650, Thorlabs, Montreal,Canada) is used to block the scattered red excitation light from themedium during fluorescence-mode imaging. Two-by-two binning of theimages enhances the signal-to-noise of the fluorescence images withreasonable acquisition times (50-1000 ms.) per spatial frequency.

The CCD detector array 11 serves as a digital camera and provides imagesto an image capture or image receiving unit 13 in image processing unit15. Within the image processing unit 15 is a memory 17 and an imageprocessor 19. Image processor 19 is configured by image processingmachine readable instructions of code 21 and structured light spatialpatterns 23 in memory 17. Image processing code 21 is configured to usea lookup table 25 of precomputed predicted fluorescence modulation forconcentrations of fluorophore located at a plurality of depths below thesurface of the medium for each spatial frequency of the plurality ofspatial frequencies.

An alternative embodiment of a system 200 configured to use qdFI isillustrated in FIG. 2 , this system 200 is integrated with a surgicalmicroscope for use in surgery. Light sources 202 feed through a spatialmodulator 204, in an embodiment a projector based on a digitalmicromirror device. Patterned light is coupled to a microscope head 206where it is focused and projected as light 208 onto a surface of tissuebeing inspected, such as brain tissue 210 beneath an opened skull 212during surgery. Light from tissue 210 is received by microscope head 206where it is focused and magnified, then enters microscope body 214 wherea beamsplitter allows portion of received light to be viewed througheyepieces 216 by a surgeon, and a portion diverted through filters 218into an electronic camera 220 where images are captured. Captured imagesfrom electronic camera 220 are fed to a digital image processing system222 through a camera interface 224. The images are processed asdescribed below with reference to FIG. 3 in an image processor 226operating under control of machine readable instructions in memory 228.Memory 228 also contains illumination light patterns, which the imageprocessor 226 may provide to spatial modulator 204 to generate spatiallymodulated light having predetermined spatial frequencies; in anembodiment the illumination light patterns include patterns with 0.1,0.3, 0.5, 0.6, 0.8 and 1.0 lines per millimeter.

Both systems of FIG. 1A and FIG. 2 operate using the method 300 of FIG.3 . To determine 302 optical parameters, including absorptioncoefficients and scattering coefficients at excitation wavelength, asurface of the medium or tissue is illuminated by projecting 304 a broadbeam of light at an excitation wavelength of the fluorophore onto thesurface using a spatially modulated pattern comprising alternating highand low intensity regions at a spatial frequency of one of thepredefined patterns. Light received from the surface is imaged 306 incamera and recorded in the image processor, these images contain apattern of diffusely reflected light emitted by the surface of themedium to form an excitation wavelength image. The steps of projectinglight and imaging are repeated 308 at at least two additional spatialfrequencies, such that all 6 available illumination light patternshaving a defined spatial frequency are used including three havingspatial frequency between 0.2 and 0.8 mm-1. Light at each spatialfrequency is projected at more than one spatial phase, effectivelyshifting the light and dark bars of the spatially modulated lightpatterns so that all portions of the surface of the medium areilluminated in at least one phase of each pattern. A model of lightpropagation in an optically absorbing and scattering medium is used tocalculate 310 an absorption coefficient and transport scatteringcoefficient of the medium at the excitation wavelength of thefluorophore at pixels of the excitation wavelength images.

Optical parameters at pixels are then determined at an emissionswavelength of the fluorophore expected to be present in inclusionswithin the medium, such as PpIX within tumors in brain tumors after ALAadministration. In an embodiment, this is done by repeating steps ofillumination at the emissions wavelength with spatially modulated lightat a plurality of phases, recording images, and extracting parametersfrom the images as done at the excitation wavelength; in an alternativebut less precise embodiment this is done by extrapolation from thestimulus wavelength parameters.

Then, while illuminating 314 with spatially modulated patterns atexcitation wavelength, fluorescent emissions images of the surface ofthe medium are recorded 316 at a plurality of spatial frequencies, inembodiments the same plurality of spatial frequencies used whileobtaining optical parameters are used by iterating 318 patternssequentially. From these images, measured modulation amplitudes areextracted 320 at the plurality of spatial frequencies for pixel blocks,such as 15×20 pixel blocks, in the fluorescent emissions images recordedat the plurality of spatial frequencies; these modulation amplitudes arethen normalized 322.

Prior to obtaining images, a lookup table 232 containing predictedfluorescence modulations for concentrations of fluorophore located at avariety of depths below the surface of the medium has been computed 324for each spatial frequency of the plurality of spatial frequencies usinga light propagation model.

The measured modulated amplitudes obtained at the plurality of spatialfrequencies for pixel regions at the plurality of spatial frequenciesare then fit 326 to the precomputed values in the lookup table todetermine depth of the concentration of fluorophore. Once the depth ofthe concentration of fluorophore is determined, the depth and at leastthe absorption parameter are used to compensate for absorption, an imagerepresenting concentration and depth may be generated, such as an imagewith brightness representing quantity of fluorophore and colorrepresenting depth.

In embodiments, the excitation wavelength is less than 100 nanometersshorter than the emissions wavelength, such as 632 nanometer stimuluswavelengths with 705 nanometer emissions wavelengths.

First Example of Operation

A system according to FIG. 1A projects six spatial patterns sequentiallyat spatial frequencies of 0, 0.05, 0.1, 0.2, 0.27 and 0.38 mm⁻¹. Theseparticular frequencies are selected to illustrate the method. Othercombinations could be used if they provide a sufficient range ofmodulation of the detected spatial light pattern. Spatially- andspectrally-resolved images are collected at each spatial frequency inboth reflectance mode and in fluorescence mode and analyzed.

Each periodic pattern of light is a spatially modulated pattern, forpurposes of this document a spatially modulated pattern is a patternhaving multiple, alternating, light and dark bars across a field ofview, the light and dark bars having a spatial frequency equal to aninverse of a pitch of the dark bars. Sine-wave spatially modulatedpatterns include patterns where, if a line is drawn perpendicular to thebars of the pattern and light intensity is plotted along the line, lightintensity varies sinusoidally along the line from a minimum at center ofdark bars to a maximum at center of light bars. In an alternativeembodiment, six patterns of structured light are projected in sequenceat spatial frequencies of 0.1, 0.3, 0.5, 0.6, 0.8, and 1 mm-1. Inembodiments, at least three spatial frequencies between 0.2 and 0.8 mm-1are used at three spatial phases for extraction of absorption andtransport scattering coefficients of the medium at both an excitationwavelength of the fluorophore PpIX and at an emissions wavelength of thefluorophore PpIX. In addition, in some embodiments a solid highintensity pattern is projected.

In one demonstration of the method and device, liquid phantoms wereprepared to simulate the optical absorption and transport scatteringcoefficients of the tissue of interest across the appropriate wavelengthrange. These phantoms included Intralipid, a lipid-protein suspension ofknown optical scattering, as well a green absorbing dye of knownextinction coefficient (absorption coefficient per unit concentration)and concentration. The Intralipid and dye concentrations were selectedto be have absorption and transport scattering coefficients at around635 and 705 nm that are typical of normal brain tissue, thesewavelengths representing, respectively, the longwave PpIX fluorescenceexcitation wavelength and PpIX emission wavelengths. A region offluorescence 2 containing a concentration of PpIX fluorophore, waspositioned at a variable distance below the surface of the phantom tosimulate, for example, a subsurface tumor. In an embodiment thisincluded a transparent cylindrical inclusion (3.2 mm inner diameter, 4.8mm outer diameter) passing through the phantom and parallel to thesurface, and filled with phosphate buffered serum containing PpIX, andeither Intralipid or blood at concentrations that yielded opticalproperties like those of the bulk medium, but with the added fluorophorePpIX. Green dye was used instead of rat blood to simulate tissueabsorption in the bulk medium, because of the limited supply of bloodavailable. The sub-surface depth of the fluorescent inclusion wasaltered by removing or adding liquid to the container, keeping a fixeddistance between the instrument and the inclusion. Phantom preparationand measurements were repeated three times for all experiments. Table 1lists the optical coefficients used in the phantom.

TABLE 1 Liquid phantom optical properties at the excitation (632 nm,‘x’) and detection (705 nm, ‘m’) wavelengths, and PpIX concentration inthe inclusion PpIX μ_(a,x) μ′_(s,x) μ_(a,m) μ′_(s,m) concentration(mm⁻¹) (mm⁻¹) (mm⁻¹) (mm⁻¹) (μg/ml) 0.019 1.17 0.001 1.0 10 0.029 1.740.0015 1.5 10 0.038 2.32 0.002 2.0 15

In a second demonstration of the method and device, minced ex vivobovine muscle was used to represent more realistic biologicalconditions. The tissue was placed in a container like that used for theliquid phantoms and a thin cylindrical inclusion (1.6 mm inner diameter,3.2 mm outer diameter) containing a fixed PpIX concentration of 5 μg/mlwas placed on top. Additional tissue then covered the inclusion to 3different depths (2.5, 4.5, 9 mm), and triplicate images were collectedat each depth. A smaller inclusion diameter was used than in the liquidphantoms to minimize air gaps between the inclusion and the tissue thatcould disturb the light pattern in the tissue.

Reflectance-mode SFDI imaging was used to map the absorption andtransport scattering properties of the phantoms. Each of the 6 spatialfrequencies, f, from 0 to 0.38 mm⁻¹ were projected onto the phantomsurface for each of 3 spatial phases, ϕ_(i)=[0, 2π/3, 4π/3]. Since theexcitation wavelength used here was 632 nm and the LCTF had limitedsensitivity above 680 nm, the optical properties around the PpIXemission wavelength band were not directly recovered. Rather, thehemoglobin absorption and Intralipid reduced scattering coefficientswere extrapolated to 705 nm. White-light SFDI was performed both at thebeginning and end of each experiment to confirm that there were nochanges in the phantom optical properties due, for example, to settlingof the Intralipid or blood.

In fluorescence mode only the 632 nm source was used, projectingsinusoidal patterns at the same spatial frequencies as above.Alternative embodiments with light sources between 620 and 640nanometers are expected to produce similar results to the experiment's632 nanometer light source. A planar fluorescence image of the inclusionin air was also captured at the beginning as well as at the end of eachexperiment to measure any signal loss due to photo-bleaching or possibleseparation of Intralipid and blood in the inclusion itself. Thecontainer was then filled with 200 ml of bulk medium and withdrawnstepwise (5 ml=1 mm change in depth) until the fluorescence from theinclusion was just detectable above background (signal/background>1.5).This was taken to represent the maximum detectable depth. Thefluorophore depth is defined here as the distance between the phantomsurface and the top of the inclusion. Planar and SFD fluorescence imageswere then acquired for each depth, integrating over 660-720 nm. Theliquid medium was removed or added randomly to vary the depth withoutintroducing systematic bias.

The measured data from the phantom were analyzed as follows to calculatethe subsurface depth, z, of the inclusion in the medium. Firstly, a 10by 20 pixel (˜2.4 mm by 4.8 mm) region of interest at the center of theinclusion was selected to determine the modulation amplitude, Mf, as afunction of f at 705 nm. Mf depends on both the fluorophore and theimaging system. Hence, at each unknown depth, Mf was normalized to theDC modulation amplitude (f=0), so that this was independent of thefluorophore concentration or quantum yield. The measured modulationamplitude was then corrected by the modulation transfer function of theimaging system, MTF_(system) (x,y,z,f). Thus, for a fluorescent objectat depth z and irradiated with unit light intensity at the fluorescenceexcitation wavelength, the normalized modulation amplitude is given byMf(x,y,f)_(z) =MTF _(system)(x,y,f)_(z) ×F _(m)(x,y,f)_(z),  (1)where F_(m)(x,y,f)_(z) is the normalized spatially-resolved diffusefluorescence in the frequency domain.

Using diffusion theory, normalized and corrected spatially-illuminatedmodulation amplitude F_(m) was calculated as a function of f for a rangeof depths z (0 to 30 mm in 0.05 mm increments) and for a range ofoptical absorption and transport scattering coefficients. A look-uptable of modulation amplitudes indexed by absorption and transportscattering coefficients and frequency was created from this set of data.

The normalized modulation amplitude in the phantom was determined as afunction of spatial frequency from the fluorescence-mode SFDImeasurements. The optical absorption and transport scatteringcoefficients of the phantom at the fluorescence emission wavelength weredetermined from the reflectance-mode SFDI measurements. Using linearinterpolation between the values in the look-up table, a least-squaresfit was performed between i) the measured modulation amplitude versusspatial frequency and ii) the corresponding predicted curve based on themeasured optical coefficients. In this fit, the fluorescent inclusiondepth was the free variable to be determined.

Knowing thereby the value of z, the PpIX concentration at this depth ineach image pixel was calculated from the measured unmodulatedfluorescence image (f=0), corrected for the effects of absorption andscattering of the fluorescence excitation and emission light by theintervening tissue. This correction was obtained by forward modeling oflight propagation in the medium, using diffusion theory with themeasured optical absorption and transport scattering coefficients asinput.

The results of these experiments in the tissue-simulating liquidphantoms and the ex vivo tissue phantom were as follows.

The modulation of the fluorescent intensity pattern measured at thesurface of one of the liquid phantoms when illuminated with 632 nm lightat different modulation frequencies, as shown in FIG. 4 , demonstrate acritical aspect of the invention, namely the fact that the rate ofdecrease of the fluorescence modulation with increasing spatialfrequency is dependent on the depth of the fluorescent inclusion belowthe surface of the optically absorbing and scattering medium.

Examples of plots of fluorescence modulation versus fluorophore depthare shown in FIG. 5A-5C and demonstrate that the plots depend on theparticular combination of absorption and transport scatteringcoefficients of the medium at the fluorescence emission wavelength. Thisthen shows that it is necessary, for example using reflectance-mode SFDIimaging, to also measure these optical coefficients so that the depth ofthe inclusion can be calculated from the measured fluorescencemodulation versus spatial frequency.

FIG. 6A-6C shows examples of the scatter plot of calculated depth versusthe actual depth of the fluorescent inclusion below the surface of thephantom. The goodness of fit, R², to the line of equality was >0.95 inall cases. These results demonstrate that the depth can be determined,within an average uncertainty of ±0.42 mm. The maximum depth that couldbe measured within this level of uncertainty ranged from 5.5 to 7.5 mmdepending on the phantom optical properties and on the PpIXconcentration. The uncertainty in the depth measurement increased to±1.5 mm at larger depths due to lower signal-to-background ratio.

FIG. 7A-7C shows examples of the corresponding graphs of PpIXconcentration in the subsurface inclusion and comparison with the actualconcentration used in the experiments. The p values indicate that thecalculated PpIX concentrations at different depths are not statisticallydifferent from the actual values in the subsurface fluorescentinclusion.

TABLE 2 Depth and PpIX concentration estimations for one phantom(μ_(a,m) = 0.002 mm⁻¹, μ′_(s,m) = 2 mm⁻¹) at an inclusion depth of 1.75mm for varying PpIX concentration PpIX 5 ± 0.007 7.5 ± 0.011 11 ± 0.01515 ± 0.02 concentration (μg/m1) Recovered 2.28 ± 0.21 1.74 ± 0.20 2.0 ±0.10 1.6 ± 0.11 depth (mm) Recovered 5.5 ± 0.6 8.2 ± 1.0 11.8 ± 0.5 15.6± 1.0 PpIX concentration (μg/ml)

The results for the ex vivo tissue model are summarized in Table 3. Themeasured tissue optical properties, for excitation at 632 nm anddetection at 705 nm, averaged over three locations near the inclusion,were μ_(ax)=0.012±0.0007, μ′_(s,x)=2.74±0.14, μ_(a,m)=0.0046±0.00014 andμ′_(s,m)=2.24±0.085 mm⁻¹. The recovered depths were within 10% of thenominal depths except at 9 mm, where the demodulated images had limitedcontrast (signal-to-background=1.8). The derived PpIX concentrationswere within 15% of actual values for the first two depths, increasing to40% at the largest depth.

TABLE 3 Summary of the estimated depth and PpIX concentration in ex vivotissue with PpIX concentration of the inclusion = 5 μg/ml. Nominal depth2.5 4.5 9 (mm) recovered depth 2.3 ± 0.02 5.0 ± 0.07 6.48 ± 0.03* (mm)recovered PpIX 4.88 ± 0.32 6.4 ± 0.85 7.04 ± 0.78* concentration (μg/ml)*indicates p < 0.05

The results from the tissue-simulating phantoms with relatively highPpIX concentration (FIG. 5 ) show that the fluorophore depth andconcentration are somewhat overestimated at large depths (7-9 mm).Conversely, the studies in ex vivo tissues (Table 3) indicate that thedepth and PpIX concentration are underestimated at large depths (9 mm)when the inclusion had a relatively low fluorophore concentration of 5μg/ml. This different behavior according to PpIX concentration was alsoseen in the subsurface depth technique of Kolste et al. Nevertheless, inall cases the errors in recovering both depth and the PpIX concentrationwere markedly reduced at higher signal-to-background (>1.5), thebackground being due, for example, to tissue autofluorescence. For aminimum signal-to-background ratio of 2, the depth was recovered towithin ±0.45 mm and the PpIX concentration within ±10% of the actualvalues at 95% confidence level for PpIX-rich tumors located from justbelow the surface up to a maximum depth of 5-9 mm, depending on thetissue turbidity. At low signal-to-background ratio the depth wasrecovered within ±1 mm, while the PpIX concentration was recoveredwithin an average accuracy of ±25%.

Further improvements of modifications to the method and device may beenvisaged. Firstly, the calibration procedure described above adds tothe steps required. Alternatively, multiple fluorescent inclusions canbe encapsulated in gels and spaced laterally at different depths forimaging simultaneously to provide the instrument calibration.

Secondly, other improvements would reduce the acquisition time andincrease the sensitivity to detect lower fluorophore concentrations atlarger depths. For example, the red LED excitation source may bereplaced by a higher power diode laser at a similar wavelength. Thevisible-range LCTF may be replaced by an LCTF with high near-infraredtransmission (for example, Varispec SNIR/NIRR, Perkin Elmer. Inc,Waltham, Mass.) or other spectral filters that are efficient in thenear-infrared range. A more sensitive imaging detector system such as anEM-CCD camera would improve the sensitivity and reduce the acquisitiontime. Direct computation of μ_(a,m) and μ′_(s,m) at all wavelengths withreflectance-mode SFDI would then be possible instead of extrapolatingthese values from 632 nm, as was done in the experiments reported above.Fast quantitative mapping of optical properties is possible with singlesnapshot SFDI, as recently reported. This approach could be incorporatedinto the present invention.

Combinations

It is anticipated that the various concepts herein disclosed can becombined in various ways. Among combinations anticipated by theinventors are:

A method designated A for determining a depth of a concentration offluorophore lying beneath a surface of an optically absorbing andscattering medium including illuminating a surface of the medium with abroad beam of light at an excitation wavelength of the fluorophore usinga first spatially modulated pattern comprising alternating high and lowintensity regions at a first spatial frequency between 0.1 and 1 mm⁻¹,the first spatially modulated pattern projected onto the surface of themedium at a plurality of different phases. While so illuminated, themethod continues with imaging diffusely reflected light emitted by thesurface of the medium to form an excitation wavelength image for eachphase of the plurality of different phases of the first spatiallymodulated pattern. The method continues with illuminating the surface ofthe medium with a broad beam of light at the excitation wavelength ofthe fluorophore using a second spatially modulated pattern comprisingalternating high and low intensity regions at a second spatial frequencybetween 0.1 and 1 mm-1, the second spatially modulated pattern projectedonto the surface of the medium at a plurality of different phases, thesecond spatially modulated pattern having a different spatial frequencythan the first spatially modulated pattern, and imaging a pattern ofdiffusely reflected light emitted by the surface of the medium to forman excitation wavelength image for each phase the plurality of differentphases of the second spatially modulated pattern. The images areprocessed by applying a model of light propagation in an opticallyabsorbing and scattering medium to calculate an absorption coefficientand transport scattering coefficient of the medium at the excitationwavelength of the fluorophore at pixels of the excitation wavelengthimages. The method also includes determining emissions wavelengthoptical parameters coefficients at pixels. The method illuminates thesurface of the medium with spatially modulated patterns at theexcitation wavelength, recording fluorescent emissions wavelength imagesof the surface of the medium at a third spatial frequency; and whileilluminating the surface of the medium with spatially modulated patternsat the excitation wavelength, recording fluorescent emissions wavelengthimages of the surface of the medium at a fourth spatial frequency. Oncethe emissions wavelength images are recorded, the method continues withextracting measured modulated amplitude at each of the third and fourthspatial frequencies for pixel regions in the fluorescent emissionswavelength images recorded at the third and fourth spatial frequencies.The method includes predicting modulated amplitude in images forconcentrations of fluorophore located at a plurality of depths below thesurface of the medium for each of the third and fourth of spatialfrequencies, and fitting the measured modulated amplitude at the thirdand fourth spatial frequencies for the pixel regions to correspondingpredicted modulated amplitudes to determine depth of the concentrationof fluorophore. The third and fourth spatial frequencies comprise atleast two spatial frequencies between 0.1 and 0.1 mm-1.

A method designated AA including the method designated A where theexcitation and emissions wavelengths are red or infrared wavelengthsgreater than or equal to 630 nanometers.

A method designated AB for determining a depth of a concentration offluorophore lying beneath a surface of an optically absorbing andscattering medium including illuminating a surface of the medium with abroad beam of light at an excitation wavelength of the fluorophore usinga first spatially modulated pattern comprising alternating high and lowintensity regions at a first spatial frequency between 0.1 and 1 mm-1,the first spatially modulated pattern projected onto the surface of themedium at a plurality of different phases. While so illuminated, themethod continues with imaging diffusely reflected light emitted by thesurface of the medium to form an excitation wavelength image for eachphase of the plurality of different phases of the first spatiallymodulated pattern. The method continues with illuminating the surface ofthe medium with a broad beam of light at the excitation wavelength ofthe fluorophore using a second spatially modulated pattern comprisingalternating high and low intensity regions at a second spatial frequencybetween 0.1 and 1 mm-1, the second spatially modulated pattern projectedonto the surface of the medium at a plurality of different phases, thesecond spatially modulated pattern having a different spatial frequencythan the first spatially modulated pattern, and imaging a pattern ofdiffusely reflected light emitted by the surface of the medium to forman excitation wavelength image for each phase the plurality of differentphases of the second spatially modulated pattern. The images areprocessed by applying a model of light propagation in an opticallyabsorbing and scattering medium to calculate an absorption coefficientand transport scattering coefficient of the medium at the excitationwavelength of the fluorophore at pixels of the excitation wavelengthimages. The method also includes determining emissions wavelengthoptical parameters coefficients at pixels. The method illuminates thesurface of the medium with spatially modulated patterns at theexcitation wavelength, recording fluorescent emissions wavelength imagesof the surface of the medium at a third spatial frequency; and whileilluminating the surface of the medium with spatially modulated patternsat the excitation wavelength, recording fluorescent emissions wavelengthimages of the surface of the medium at a fourth spatial frequency. Oncethe emissions wavelength images are recorded, the method continues withextracting measured modulated amplitude at each of the third and fourthspatial frequencies for pixel regions in the fluorescent emissionswavelength images recorded at the third and fourth spatial frequencies.The method uses the emissions wavelength images to determine depth ofthe concentration of fluorophore. The third and fourth spatialfrequencies comprise at least two spatial frequencies between 0.1 and0.1 mm-1, and the excitation wavelength and emissions wavelength areboth in the “red” portion of the spectrum traditionally known as 630-740nanometer wavelengths, or in the infrared portion of the spectrum withwavelengths 740 nanometers and longer.

A method designated AC including the method designated A, AA, or ABwherein the spatially modulated patterns comprise a plurality ofalternating light and dark bars.

A method designated AD including the method designated A, AA, AB, or ACwherein determining emissions wavelength optical parameters at pixels inthe images is performed by a method including irradiating a surface ofthe medium with a broad beam of light at the emissions wavelength of thefluorophore using a spatially illuminated pattern comprising alternatinghigh and low intensity regions at a fifth spatial frequency between 0.1and 1 lines per millimeter at each of a plurality of phases; imaging apattern of diffusely reflected light emitted by the surface of themedium to form an emissions-wavelength, spatially-illuminated, image forthe fifth spatial frequency at each of the plurality of phases;irradiating a surface of the medium with a broad beam of light at theemissions wavelength of the fluorophore using a spatially illuminatedpattern comprising alternating high and low intensity regions at a sixthspatial frequency between 0.1 and 1 lines per millimeter at each of aplurality of phases; imaging a pattern of diffusely reflected lightemitted by the surface of the medium to form an emissions-wavelength,spatially-illuminated, image for the sixth spatial frequency at each ofthe plurality of phases; and applying a model of light propagation in anoptically absorbing and scattering medium to calculate an absorptioncoefficient and transport scattering coefficient of the medium at theemissions wavelength of the fluorophore at each pixel in the images.

A method designated AE including the method designated A, AA, AB, AC, orAD and further including using the depth of the concentration offluorophore and at least the absorption coefficient at each pixel toquantify the concentration of fluorophore.

A method designated AF including the method designated A, AA, AB, AC,AD, or AE in which the fluorescent material is protoporphyrin IX.

A method designated AG including the method designated A, AA, AB, AC,AD, AE, or AF further comprising preparing an image with fluorophoreconcentration encoded as brightness and fluorophore depth encoded ascolor.

A method designated AH including the method designated AF in which theabsorbing and scattering medium is mammalian tissue and theprotoporphyrin IX is generated endogenously in the tissue uponadministration of 5-aminolevulinic acid.

A method designated AJ including the method designated A, AA, AB, AC,AD, AE, AF, AG, or AH in which the concentration of fluorophorecomprises a tumor.

A method designated AK including the method designated A, AA, AB, AC,AD, AE, AF, AG, or AH in which the concentration of fluorophorecomprises a molecular material or nanomaterial that fluoresces in thered or near-infrared spectral range from about 600 nm to 1300 nm.

A method designated AL including the method designated AD, AE, AF, AG,AH, AJ, or AK wherein the first, third, and fifth spatial frequenciesare the same, the second, fourth, and sixth spatial frequencies are thesame, and first and second spatial frequencies differ.

A device designated B for quantifying and determining depth ofconcentrations of fluorophores includes: an excitation-wavelength lightsource configurable to excite the fluorophores in the medium using aspatially-modulated light pattern, the spatially-modulated light patternselectable from a plurality of light patterns with spatial frequenciesbetween 0.1 and 1 lines per millimeter; at least one optical elementconfigured to transfer diffusely-reflected images and fluorescenceemission images from the medium to an imaging subsystem; the imagingsubsystem comprising at least one filter insertable into the light pathto reject light at the excitation wavelength, and at least oneelectronic camera configured to convert received light into a matrix ofdigital values according to an intensity of light striking detectorpixels; and at least one computer configured to control thespatially-modulated light pattern, to capture images, and configuredwith machine readable instructions to determine a depth of theconcentration of fluorophore and determine a concentration of thefluorophore for image pixels of the camera; and apparatus configured todisplay fluorophore concentration depth and concentration.

A device designated BA including the device designated B wherein theexcitation wavelength is within 100 nanometers of a wavelength peak offluorescent emissions of the concentration of fluorophore.

A device designated BB including the device designated B or BA whereinthe machine readable instructions to determine a depth of aconcentration of fluorophore are configured for a method including:extracting measured modulated amplitude at the plurality of spatialfrequencies for pixel regions in fluorescent emissions images recordedat a plurality of spatial frequencies; predicting fluorescencemodulation for a concentration of fluorophore located at a plurality ofdepths below the surface of the medium for each spatial frequency of theplurality of spatial frequencies; and fitting the measured modulatedamplitude at the plurality of spatial frequencies for pixel regions atthe plurality of spatial frequencies to corresponding predictedfluorescence modulation to determine depth of the concentration offluorophore.

A device designated BC including the device designated B, BA, or BB inwhich the absorbing and scattering medium is mammalian tissue.

A device designated BD including the device designated B, BA, BB, or BCwherein the excitation wavelength has wavelength between 620 and 640nanometers.

A device designated BE including the device designated B, BA, BC, or BDwherein the fluorescent emissions wavelength has a peak between 700 and720 nanometers, as does PpIX.

Changes may be made in the above methods and systems without departingfrom the scope hereof. It should thus be noted that the matter containedin the above description or shown in the accompanying drawings should beinterpreted as illustrative and not in a limiting sense. The followingclaims are intended to cover all generic and specific features describedherein, as well as all statements of the scope of the present method andsystem, which, as a matter of language, might be said to falltherebetween.

We claim:
 1. A device for quantifying and determining depth of concentrations of fluorophores in a turbid medium comprising: an excitation-wavelength light source configurable to excite the fluorophores in the medium using a spatially-modulated light pattern, the spatially-modulated light pattern selectable from a plurality of light patterns with spatial frequencies between 0.2 and 1 lines per millimeter; an optical element configured to transfer light from the turbid medium to an imaging subsystem; the imaging subsystem comprising at least one filter insertable into a light path from the turbid medium to an electronic camera to reject light at the excitation wavelength; and at least one computer configured to control the spatially-modulated light pattern, to capture diffusely-reflected images and fluorescence emission images using the electronic camera, and configured with machine readable instructions to determine a depth of the concentration of fluorophore and determine a concentration of the fluorophore for pixels of the diffusely-reflected and fluorescence emission images; and apparatus configured to display the fluorophore concentration depth and concentration.
 2. The device of claim 1 wherein the machine readable instructions to determine a depth of the concentration of fluorophore are configured for a method comprising: extracting measured modulated amplitude at the plurality of spatial frequencies for pixel regions in the fluorescent emissions images recorded at a plurality of spatial frequencies; predicting fluorescence modulation for a concentration of fluorophore located at a plurality of depths below the surface of the medium for each spatial frequency of the plurality of spatial frequencies; and fitting the measured modulated amplitude at the plurality of spatial frequencies for pixel regions at the plurality of spatial frequencies to corresponding predicted fluorescence modulation to determine depth of the concentration of fluorophore; and wherein the excitation wavelength is within 100 nanometers of a wavelength peak of fluorescent emissions of the concentration of fluorophore.
 3. The device of claim 2 wherein the machine readable instructions comprise instructions for determining optical properties of the turbid medium at both the excitation wavelength and the emissions wavelength.
 4. The device of claim 2 wherein the camera is a spectrally resolved camera adapted to capture separate images at the fluorescent stimulus and the fluorescent emissions wavelengths.
 5. The device of claim 2 in which the turbid medium is mammalian tissue.
 6. The device of claim 5 wherein the excitation wavelength is between 620 and 640 nanometers.
 7. The device of claim 6 wherein the fluorescent emissions wavelength has a peak lying between 700 and 720 nanometers.
 8. The device of claim 5, wherein the fluorescent emissions wavelength has a peak lying between 700 and 720 nanometers.
 9. The device of claim 8 wherein the spectrally resolved camera is adapted to capture separate images at a first and a second fluorescent emissions wavelength, and the machine readable instructions include instructions for computing ratios of emissions at the first and second fluorescent emissions wavelengths.
 10. The device of claim 9 wherein the fluorescent stimulus wavelength is red or infrared.
 11. The device of claim 10 configured to resolve fluorophore concentrations at depths up to nine millimeters in tissue.
 12. The device of claim 2 further comprising a phantom having fluorophore concentrations at a plurality of known depths.
 13. A device for quantifying and determining depth of concentrations of fluorophores in a turbid medium comprising: an excitation-wavelength light source configurable to excite the fluorophores in the medium using a spatially-modulated light pattern, the spatially-modulated light pattern selectable from a plurality of light patterns with spatial frequencies between 0.2 and 1 lines per millimeter; an optical element configured to transfer light from the turbid medium to an imaging subsystem; the imaging subsystem being a spectrally resolved imaging system comprising at least one tunable filter in a light path from the turbid medium to an electronic camera; and at least one computer configured to control the spatially-modulated light pattern, to capture diffusely-reflected images and fluorescence emission images using the electronic camera, and configured with machine readable instructions to determine a depth of the concentration of fluorophore and determine a concentration of the fluorophore for pixels of the diffusely-reflected and fluorescence emission images; and apparatus configured to display the fluorophore concentration depth and concentration.
 14. The device of claim 13 wherein the machine readable instructions are configured to compute optical parameters of the turbid medium at both excitation and emissions wavelengths.
 15. The device of claim 14 wherein: the spectrally resolved imaging system is adapted to capture separate images at a first and a second fluorescent emissions wavelength, the machine readable instructions include instructions for computing ratios of emissions at pixels of the images at the first and second fluorescent emissions wavelengths, and the machine readable instructions for determining depth of the concentration of fluorophore use the ratios of emissions at the first and second fluorescent emissions wavelengths in determining depth of the concentration of fluorophore.
 16. The device of claim 15 wherein the machine readable instructions to determine a depth of a concentration of fluorophore are configured for a method comprising: extracting measured modulated amplitude at the plurality of spatial frequencies for pixel regions in fluorescent emissions images recorded at a plurality of spatial frequencies; predicting fluorescence modulation for a concentration of fluorophore located at a plurality of depths below the surface of the medium for each spatial frequency of the plurality of spatial frequencies from the ratios of emissions at the first and second fluorescent emissions wavelengths; and fitting the measured modulated amplitude at the plurality of spatial frequencies for pixel regions at the plurality of spatial frequencies to corresponding predicted fluorescence modulation to determine depth and concentration of the concentration of fluorophore.
 17. The device of claim 16 wherein the machine readable instructions comprise instructions to correct concentration of the concentration of fluorophore using the computed optical parameters of the turbid medium at the fluorescent stimulus and emissions wavelengths.
 18. The device of claim 17 wherein the device is capable of resolving concentrations of fluorophore at depths up to nine millimeters in tissue. 